For many years, enzyme assays and immunoassays have been successfully used in a wide variety of clinical, veterinary, and bio-analytical laboratory applications. Increasingly, as the availability of monoclonal and polyclonal antibodies has expanded, the range of applications to which these antibodies has extended include environmental monitoring applications (Vanderlaan et. al., 1988). Enzyme and antibody assay technology ahs been used to measure drug abuse, hormones, to monitor therapeutic drugs and to screen for environmental pollutants. These tests are generally simple to use, sensitive, and inexpensive. they are typically, however, colorimetric assays and are at best semi-quantative.
Biosensors offer an alternative format for the performance of established enzyme assays and immunoassays and create new opportunities for the application of established assays to continuous monitoring situations. Biosensors also offer distinct advantages (Taylor, 1987) in being very rapid (sub-second response rates), with condiderably improved limits of detection (typically ppb), are quantitative, and provide electronic outputs that may be integrated into simple data loggers or complex feedback control systems.
Parallel significant advances in microelectronics fabrication technology and biotechnology have created a unique interface of materials science and biology. Biosensors are measurement devices that detect and discriminate among molecules or substances of interest (analytes) through the recognition reactions of biologically active molecules (Lowe, 1985). The biosensor is a microelectronic device or chip, while biorecognition molecules are bioactive proteins. The biosensor measures the concentration of an analyte and produces a proportionate, electrical signal. In biosensors, biologically active molecules provide the needed molecular specificity and confer the ability to detect and discriminate among various substances to be analyzed. Microfabricated solid state devices provide the means for electrical transduction. Together and through their association, these two elements combine to produce an electrically based measurement signal that is the result of the physical chemical changes associated with biorecognition and serve as the raw output data of the biosensor test (Thompson and Krull, 1991).
FIG. 1 schematically illustrates the layout for a typical biosensor instrument. Typically, the biosensor instrument consists of three functionally distinct parts; the biotransducer or biosensor device, the device interrogator and signal processor, and the output device.
Biosensor devices are typically configured as shown in the schematic of FIG. 2. Biosensor devices typically contain two basic parts. The first part is an organic biorecognition layer that contains the biorecognition molecules. The second part is a microdevice which provides the electrical signal and is an electrical or electronically based component. Together, these components form the active element or biotransducer of the biosensor instrument. The biologically active component may be a thin organic film of enzyme, immunochemically active protein, stabilized receptor, a tissue slice, or a cell fragment. This part provides the molecular specificity o recognition of the analyte present in the test sample or process stream. The biologically active part of the transducer converts the analyte into another substance or produces some physical chemical change which can be detected by the electrically or electronically active part of the biotransducer. The electronically or electrically active part of the biotransducer accordingly produces a current or voltage change in response to the appearance of the product of the biological transmutation or conversion reaction. The current or voltage change is an analog signal which may be amplified locally before going on to the second part of the biosensor instrument--the signal processor.
Biosensor devices are integral components of biosensor instruments. The signal processor captures the signal from the biotransducer, may amplify, smooth or perform some mathematical operations on the data, then present it to the next part of the instrument--the output. The output section is responsible for presenting the acquired data in a form suitable for the senses and compatible with other information sources to produce sound and timely decision making by the end user.
This general description of biosensor technology encompasses a wide range of bioanalytical devices, some of which are pH or ion selective electrodes, while others are complicated optical devices.
Electroactive polymer biosensor devices must transmute the chemical potential energy associated with the concentration of an analyte into a proportionate, measurable, electrical signal. The specificity of these electroactive polymer biosensor devices is derived for the biorecognition reactions of immobilized enzymes, enzyme-linked antibody conjugates, and stabilized receptors. Several approaches to the use of electroactive polymers are possible. These are potentiometric, amperometric, chemoresistive, and methods based upon field effect phenomena. In all these methods the electroactive polymer serves as an active, functional, and integral part of the bio-transducer.
Potentiometric biosensor devices derive their responses from the changes in the steady state potential which accompanies the change in redox composition of the electroactive polymer film. This is a direct detection method not requiring the input of externally generated energy. Films may be free standing or supported by ohmic contact to a inert electrode. Changes in extent of oxidation are reflected in changes in the population of polaron o bipoloran states and accompanying counterion ingress. Thus electroactive polymer films therefore serve as ionophore containing membranes of ion-selective electrodes (ISEs). The membrane potential of the electroactive polymer is directly related to some potentiometrically measurable ionic product of the biorecognition reaction (Thompson et al., 1986). In general, all electroactive polymers display som measurable change in steady state or open circuit potential as a function of doping level. This is true over some specified range of redox composition.
Band-gap, p-type semiconducting polymers such as polyacetylene and polypyrrole can change their open-circuit electrode potentials through a maximum of the difference between the mid-gap energy (intrinsic semiconductor) and the conduction band edge, and typically, this change is between the mid-gap energy and the Fermi energy, E.sub.F -E.sub.CB (Morrison, 1980). For polyacetylene, the open circuit potential is an invariant 0.45V vs SCE for CH.sub.x compositions above the insulator to semiconductor/metal transition (ca. 4 mol % of dopant per --C.dbd.C--unit) (Guiseppi-Elie and Wnek, 1990). Thus the maximum voltage change (over all dopant compositions) associated with these materials is 1/2 the band gap, E.sub.g, or ca. 0.70 V. Direct potentiometric detection of nucleotides co-deposited in electropolymerized polypyrrole has been reported by Shimidzu (1987) and potentiometric measurements of the concentrations of various anions by electropolymerized polypyrrole films have been reported by Dong et. al. (1988). Both these works suggest interesting and unusual behavior of electroactive polypyrrole in potentiometric measurement mode. Since n-type semiconducting band-gap polymers are violently reactive in aqueous environments, their applicability to biosensors is necessarily without consideration.
Redox polymers such as poly(vinylferrocene) and poly(viologen) establish an open circuit potential which reflects the equilibria between oxidized and reduced forms of the redox-active moiety (Lewis et. al., 1984). These devices work by indirectly linking the biorecognition reaction to a change in redox composition of the film. Any signature redox active species which can alter the equilibrium at the polymer modified electrode surface provides a means for biotransduction. These systems all show behavior that can be predicted by the Nernst equation and electrochemical kinetics. Because potentiometric devices based on electroactive polymers are steady state devices they are subject to limitations arising from errors such as interfering ions, non-specific protein binding (Collins and Janata, 1982), surface charge adsorption, and the sensitivity limitations imposed by Band theory and the 59 mV/decade sensitivity of the Nernst equation (Thompson and Krull, 1991).
Amperometric biosensors formed from electroactive polymers ar by far the most common generic approach to biosensors (Janata and Bezegh, 1988). With such devices, two principles dominate: i) redox mediation and ii) electrocatalysis. Redox mediation implies the oxidoreduction of an electroactive species other than the on of interest. Umana and Waller (1986) used the amperometric reduction of iodine (I.sub.2 +2e.fwdarw.2I) mediated through the Mo.sup.IV -catalyzed reaction of H.sub.2 O.sub.2 with iodide (H.sub.2 O.sub.2 +2H.sup.+ 2I+2I.fwdarw.I.sub.2 +2H.sub.2 O) under aerobic conditions.l Iwakura et. al. (1988) achieved direct redox mediation with ferrocenecarboxylic acid under anaerobic conditions. Similar devices have since also been demonstrated by Caglar and Wnek (1991). In such devices the electroactive polymer functions simply as a retaining matrix for the mediator and is itself not actively involved in transduction. The related works of Heller (1990) and Hale et. al. (1991) use oligomers and polymers of ferrocene-based redox mediators as active participants to transduction (Gorton et. al., 1990). These are true transducer-active polymer films as they both mediate and directly transmute the redox activity of the enzyme to amperometrically discharged current at metallic or carbon paste electrode.
The simplest amperometric biosensors of this type are those based on the discharge of a redox active bi-products of the biochemical transmutation reaction. The early glucose biosensors which amperometrically discharged H.sub.2 O.sub.2 are of this type (Clark, 1987). There is evidence that electroactive polyaniline modified inert electrodes can electrocatalyze the reduction of ascorbic acid (Hepel, 1990) and hydrogen peroxide (Guiseppi-Elie and Wilson, 1990), both of which are associated with biochemical reactions. Polyaniline modified platinum electrodes have also been reported to electrocatalyze the oxidation of formic acid (Gholamian et. al., 1987). The potential for using electroactive polymer modified electrodes for electrocatalysis in Clark-typ amperometric biosensors is evident but little explored.
Chemoresistive biosensor devices based on electroactive polymer membrane films take advantage of the very large and rapid change in electrical conductivity which accompanies "doping" of the electroactive polymer (Baughman and Shacklette, 1990). The principle of operation in this approach is based on the measurement of biochemically modulated changes in the electronic resistance of the membrane film. The membrane films, fabricated on solid state devices, change their electrical impedance characteristics in response to the biological reactions with which they are associated. Chemoresistive biosensor devices are composed of a fully contiguous membrane film of electroactive polymer fabricated over the interdigit area of Interdigitated Microsensor Electrodes (IMEs) or array microelectrodes. IMEs generally have digit dimensions and separation distances which range from 1 to 20 microns. These dimensions are readily achievable with commonly available microlithography technology. The number of digits and the digit length together define the meander length for the device and this is selected to match the resistance range of the particular electroactive polymer being employed. The short separation distance and long meander length found on IMEs allow the generation of high electric fields with modest voltages while allowing effective work with highly resistive membrane materials. The device sensitivity is provided by the large (up to 12 orders of magnitude) (Frommer and Chance, 1986), dynamic ( ms - .mu.s response) (Thackeray et. al., 1985) chemoresistance range of chemically sensitive, electroactive polymer membrane films.
Chemoresistive sensor devices based on electroactive polymer films were introduced by Paul et. al. (1985). These redox dependant chemoresistance devices emphasized voltage modulation of the chemoresistive response. Thackeray et. al. (1985) describe a device based on poly(3-methylthiophene) which illustrates a key strength of this method--The low detection limit (&lt;10.sup.-15 moles of oxidant) to elicit a response above noise level and the significant amplification possible with proper design of IME and fabrication of the chemically sensitive film (Lofton et. al., 1986). While the foregoing illustrates the general principles of chemoresistance detection the electroactive polymer films used were not conferred with biospecificity and are accordingly not biosensors. Taylor et al. (1988) report receptor-based biosensors using the chemoresistance principle but the polymers used were not electroactive (Taylor, 1989). Malmros et. al. (1987/88) describe a chemoresistance biosensor based on free-standing polyacetylene. This report suggests that the measured chemoresistance response was due primarily to changes in ionic resistance attendant to increases in the extent of wetting of Shirakawa polyacetylene upo doping with aqueous H.sub.2 O.sub.2 or I.sub.2. Chemoresistance biosensors have also been reported by Watson et. al. (1987/88) and Cullen et. al. (1990).
These films are well known to bridge substantial device scale distances (Focke et. al. 1989). Using established methods designed to confer general chemical and biospecificity to these films (Guiseppi-Elie, 1988) and using these devices in a kinetic mode (Karube, 1987) opens up additional possibilities for FETs (Garnier et. al., 1991).
While there have been a large number of prior art biosensors, it would be desirable to have a system and method which can be utilized for the interrogation capture and analysis of chemoresistive sensor responses.
It is thus an object of the present invention to provide a novel analytical method which can be utilized for the interrgogation, capture and analysis of chemoresistive sensor responses.
It is a further object of the present invention to provide novel components which together comprise an electroactive polymer sensor interrogation system.
It is still yet a further object of the present invention to provide chemical and biosensor devices formed from chemically modified and derivatized electroactive polymer films
It is still an additional object of the present invention to provide an analytical method for monitoring the time rate of change (kinetic) or extent of change (equilibrium) of the resistance of the electroactive polymer film as it spontaneously reacts with a redox active analyte to which it has been rendered specific
These and other objects of the present invention will become apparent from the following summary and detailed description which follow.